Medical devices with long term non-thrombogenic coatings

ABSTRACT

A coating and method for implantable open lattice metallic stent prostheses are disclosed. The coating includes a relatively thin layer of biostable elastomeric material containing an amount of biologically active material, particularly heparin, dispersed in the coating in combination with a non-thrombogenic surface. In one embodiment, the surface is provided with sites of high electronegativity species by coating with fluorosilicone which aid in controlling elution, particularly the initial release rate, and reduced thrombogenic activity. Other non-thrombogenic outer layers for heparin such as covalently bound polyethylene glycol (PEG) are also disclosed.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is a Continuation-In-Part of application Ser.No. 08/526,273, filed Sep. 11, 1995, now abandoned, and aContinuation-In-Part of application Ser. No. 08/424,884, filed Apr. 19,1995, now abandoned, all portions of the parent applications notcontained in this application being deemed incorporated by reference forany purpose. Cross-reference is also made to Ser. No. 08/663,490,entitled "DRUG RELEASE STENT COATING PROCESS, filed of even date, ofcommon inventorship and assignee, now U.S. Pat. No. 5,837,313 and also aContinuation-In-Part of both above-referenced applications. To theextent that it is not contained herein, that application is also deemedincorporated herein by reference for any purpose.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates generally to providing biostableelastomeric coatings on the surfaces of implants which incorporatebiologically active species having controlled release characteristics inthe coating particularly to providing a non-thrombogenic surface duringand after timed release of the biologically active species. Theinvention is particularly described in terms of coatings on therapeuticexpandable stent prostheses for implantation in body lumens, e.g.,vascular implantation.

2. Related Art

In surgical or other related invasive procedures, the insertion andexpansion of stent devices in blood vessels, urinary tracts or otherlocations difficult to otherwise access for the purpose of preventingrestenosis, providing vessel or lumen wall support or reinforcement andfor other therapeutic or restorative functions has become a common formof long-term treatment. Typically, such prostheses are applied to alocation of interest utilizing a vascular catheter, or similartransluminal device, to carry the stent to the location of interestwhere it is thereafter released to expand or be expanded in situ. Thesedevices are generally designed as permanent implants which may becomeincorporated in the vascular or other tissue which they contact atimplantation.

One type of self-expanding stent has a flexible tubular body formed ofseveral individual flexible thread elements each of which extends in ahelix configuration with the centerline of the body serving as a commonaxis. The elements are wound in the same direction but are displacedaxially relative to each other and meet, under crossing, a like numberof elements also so axially displaced, but having the opposite directionof winding. This configuration provides a resilient braided tubularstructure which assumes stable dimensions upon relaxation. Axial tensionproduces elongation and corresponding diameter contraction that allowsthe stent to be mounted on a catheter device and conveyed through thevascular system as a narrow elongated device. Once tension is relaxed insitu, the device at least substantially reverts to its original shape.Prostheses of the class including a braided flexible tubular body areillustrated and described in U.S. Pat. Nos. 4,655,771 and 4,954,126 toWallsten and U.S. Pat. No. 5,061,275 to Wallsten et al.

Implanted stents have been used to carry medicinal agents, such asthrombolytic agents. U.S. Pat. No. 5,163,952 to Froix discloses athermal memoried expanding plastic stent device formulated to carry amedicinal agent in the material of the stent itself. Pinchuk, in U.S.Pat. No. 5,092,877, discloses a stent of a polymeric material which mayhave a coating associated with the delivery of drugs. Other patentswhich are directed to devices of the class utilizing bio-degradable orbio-sorbable polymers include Tang et al, U.S. Pat. No. 4,916,193, andMacGregor, U.S. Pat. No. 4,994,071.

A patent to Sahatjian, U.S. Pat. No. 5,304,121, discloses a coatingapplied to a stent consisting of a hydrogel polymer and a preselecteddrug such as cell growth inhibitors or heparin. A further method ofmaking a coated intravascular stent carrying a therapeutic material isdescribed in Berg et al., U.S. Pat. No. 5,464,650, issued on Nov. 7,1995 and corresponding to European Patent Application No. 0 623 354 A1published Nov. 9, 1994. In that disclosure, a polymer coating materialis dissolved in a solvent and the therapeutic material dispersed in thesolvent; the solvent evaporated after application.

An article by Michael N. Helmus (a co-inventor of the present invention)entitled "Medical Device Design--A Systems Approach: Central VenousCatheters", 22nd International Society for the Advancement of Materialand Process Engineering Technical Conference (1990) relates topolymer/drug/membrane systems for releasing heparin. Thosepolymer/drug/membrane systems require two distinct types of layers tofunction.

It has been recognized that contacting blood with the surface of aforeign body in vivo has a tendency to induce thrombogenic responses andthat as the surface area of a foreign device in contact with host bloodincreases, the tendency for coagulation and clot forming at thesesurfaces also increases. This has led to the use of immobilized systemicanti-coagulant or thrombolytic agents such as heparin on bloodcontacting surfaces such as oxygen uptake devices to reduce thisphenomenon. Such an approach is described by Winters, et al., in U.S.Pat. Nos. 5,182,317; 5,262,451 and 5,338,770 in which the aminefunctional groups of the active material are covalently bonded usingpolyethylene oxide (PEO) on a siloxane surface.

Another approach is described in U.S. Pat. No. 4,613,665 to Larm inwhich heparin is chemically covalently bound to plastic surfacematerials containing primary amino groups to impart a non-thrombogenicsurface to the material. Other approaches for bonding heparin aredescribed in Barbucci, et al., "Coating of commercially availablematerials with a new heparinizable material", Journal of BiomedicalMaterials Research, Vol 25, 1259-1274 (1991); Hubbell, J. A.,"Pharmacologic Modification of Materials", Cardiovascular Pathology, Vol2, No 3(Suppl.), 121S-127S (1993); Gravlee, G. P., "Heparin-CoatedCardiopulmonary Bypass Circuits", Journal of Cardiothoracic and VascularAnesthesia, Vol 8, No 2, pp 213-222 (1994).

Although polymeric stents are effective, they, may have mechanicalproperties that are inferior to those of metal stents of like thicknessand weave. Metallic vascular stents braided of even relatively finemetal can provide a large amount of strength to resist inwardly directedcircumferential pressure. A polymer material of comparable strengthrequires a much thicker-walled structure or heavier, denser filamentweave, which in turn, reduces the cross-sectional area available forflow through the stent and/or reduces the relative amount of open spacein the weave. Also, it is usually more difficult to load and deliverpolymeric stents using catheter delivery systems.

While certain types of stents such as braided metal stents may bepreferred for some applications, the coating and coating modificationprocess of the present invention is not so limited and can be used on awide variety of prosthetic devices. Thus, in the case of stents, thepresent invention also applies, for example, to the class of stents thatare not self-expanding including those which can be expanded, forinstance, with a balloon; and is applicable to polymeric stents of allkinds. Other medical devices that can benefit from the present inventioninclude blood exchanging devices, vascular access ports, central venuscatheters, cardiovascular catheters, extracorpeal circuits, vasculargrafts, pumps, heart valves, and cardiovascular sutures, to name a few.Regardless of detailed embodiments, applicability of the inventionshould not be considered limited with respect to implant design, implantlocation or materials of construction. Further, the present inventionmay be used with other types of implantable prostheses.

Accordingly, it is a primary object of the present invention to providea coating and process for coating a stent to be used as a deployed stentprostheses, the coating being capable of effective controlled long-termdelivery of biologically active materials.

Another object of the invention is to provide a coating and process forcoating a stent prostheses using a biostable hydrophobic elastomer inwhich biologically active species are incorporated within a coating.

Still another object of the present invention is to provide amulti-layer coating and process for the delivery of biologically activespecies in which the percentage of active material can vary from layerto layer.

Yet another object of the present invention is to provide a multi-layercoating and process for the delivery of biologically active species froma coating with a non-thrombogenic surface.

A further object of the invention is to provide a multi-layer coatingfor the delivery of biologically active species such as heparin having afluorosilicone top layer.

A still further object of the invention is to provide a multi-layercoating for the delivery of biologically active species such as heparinhaving a surface containing immobilized polyethylene glycol (PEG).

Other objects and advantages of the present invention will becomeapparent to those skilled in the art upon familiarization with thespecification and appended claims.

SUMMARY OF THE INVENTION

The present invention provides a relatively thin layered coating ofbiostable elastomeric material containing an amount of biologicallyactive material dispersed therein in combination with a non-thrombogenicsurface that is useful for coating the surfaces of prostheses such asdeployable stents.

The preferred stent to be coated is a self-expanding, open-ended tubularstent prostheses. Although other materials, including polymer materials,can be used, in the preferred embodiment, the tubular body is formed ofa self-expanding open braid of fine single or polyfilament metal wirewhich flexes without collapsing, readily axially deforms to an elongateshape for transluminal insertion via a vascular catheter and resilientlyexpands toward predetermined stable dimensions upon removal in situ.

In the process, the initial coating is preferably applied as a mixture,solution or suspension of polymeric material and finely dividedbiologically active species dispersed in an organic vehicle or asolution or partial solution of such species in a solvent or vehicle forthe polymer and/or biologically active species. For the purpose of thisapplication, the term "finely divided" means any type or size ofincluded material from dissolved molecules through suspensions, colloidsand particulate mixtures. The active material is dispersed in a carriermaterial which may be the polymer, a solvent, or both. The coating ispreferably applied as a plurality of relatively thin layers sequentiallyapplied in relatively rapid sequence and is preferably applied with thestent in a radially expanded state.

In many applications the layered coating is referred to or characterizedas including an undercoat and topcoat. The coating thickness ratio ofthe topcoat to undercoat may vary with the desired effect and/or theelution system. Typically these are of different formulations with mostor all of the active material being contained in the undercoat and anon-thrombogenic surface is found in the topcoat.

The coating may be applied by dipping or spraying using evaporativesolvent materials of relatively high vapor pressure to produce thedesired viscosity and quickly establish coating layer thicknesses. Thepreferred process is predicated on reciprocally spray coating a rotatingradially expanded stent employing an air brush device. The coatingprocess enables the material to adherently conform to and cover theentire surface of the filaments of the open structure of the stent butin a manner such that the open lattice nature of the structure of thebraid or other pattern is preserved in the coated device.

The coating is exposed to room temperature ventilation for apredetermined time (possibly one hour or more) for solvent vehicleevaporation. In the case of certain undercoat materials, thereafter thepolymer material is cured at room temperature or elevated temperatures.Curing is defined as the process of converting the elastomeric orpolymeric material into the finished or useful state by the applicationof heat and/or chemical agents which induce physico-chemical changes.Where, for example, polyurethane thermoplastic elastomers are used as anundercoat material, solvent evaporation can occur at room temperaturerendering the undercoat useful for controlled drug release withoutfurther curing.

The applicable ventilation time and temperature for cure are determinedby the particular polymer involved and particular drugs used. Forexample, silicone or polysiloxane materials (such aspolydimethylsiloxane) have been used successfully. Urethane pre-polymerscan also be utilized. Unlike the polyurethane thermoplastic elastomers,some of these materials are applied as pre-polymers in the coatingcomposition and must thereafter be heat cured. The preferred siliconespecies have relatively low cure temperatures and are known as a roomtemperature vulcanizable (RTV) materials. Some polydimethylsiloxanematerials can be cured, for example, by exposure to air at about 90° C.for a period of time such as 16 hours. A curing step may be implementedboth after application of the undercoat or a certain number of lowerlayers and the top layers or a single curing step used after coating iscompleted.

The coated stents may thereafter be subjected to a postcure processwhich includes an inert gas plasma treatment, and sterilization whichmay include gamma radiation, ETO treatment, electron beam or steamtreatment.

In the plasma treatment, unconstrained coated stents are placed in areactor chamber and the system is purged with nitrogen and a vacuumapplied to 20-50 mTorr. Thereafter, inert gas (argon, helium or mixtureof them) is admitted to the reaction chamber for the plasma treatment.One method uses argon (Ar) gas, operating at a power range from 200 to400 watts, a flow rate of 150-650 standard ml per minute, which isequivalent to about 100-450 mTorr, and an exposure time from 30 secondsto about 5 minutes. The stents can be removed immediately after theplasma treatment or remain in the argon atmosphere for an additionalperiod of time, typically five minutes.

In accordance with the invention, the top coat or surface coating may beapplied in any of several ways to further control thrombolitic effectsand optionally, control the release profile especially the initial veryhigh release rate associated with the elution of heparin.

In one embodiment, an outer layer of fluorosilicone (FSi) is applied tothe undercoat as a topcoat. The outer layer can also contain heparin. Inanother embodiment, polyethylene glycol (PEG) is immobilized on thesurface of the coating. In this process, the underlayer is subjected toinert gas plasma treatment and immediately thereafter is treated byammonia (NH₃) plasma to aminate the surface. Amination, as used in thisapplication, means creating mostly imino groups and other nitrocontaining species on the surface. This is followed by immediateimmersion into electrophillically activated polyethylene glycol(PEG)solution with a reductive agent, i.e., sodium cyanoborohydride.

The coated and cured stents having the modified outer layer or surfaceare subjected to a final gamma radiation sterilization nominally at2.5-3.5 Mrad. Argon (Ar) plasma treated stents enjoy full resiliencyafter radiation whether exposed in a constrained or non-constrainedstatus, while constrained stents subjected to gamma sterilizationwithout Ar plasma pretreatment lose resiliency and do not recover at asufficient or appropriate rate.

The elastomeric materials that form the stent coating underlayers shouldpossess certain properties. Preferably the layers should be of suitablehydrophobic biostable elastomeric materials which do not degrade.Surface layer material should minimize tissue rejection and tissueinflammation and permit encapsulation by tissue adjacent the stentimplantation site. Exposed material is designed to reduce clottingtendencies in blood contacted and the surface is preferably modifiedaccordingly. Thus, underlayers of the above materials are preferablyprovided with a fluorosilicone outer coating layer which may or may notcontain imbedded bioactive material, such as heparin. Alternatively, theouter coating may consist essentially of polyethylene glycol (PEG),polysaccharides, phospholipids, or combinations of the foregoing.

Polymers generally suitable for the undercoats or underlayers includesilicones (e.g., polysiloxanes and substituted polysiloxanes),polyurethanes, thermoplastic elastomers in general, ethylene vinylacetate copolymers, polyolefin elastomers, polyamide elastomers, andEPDM rubbers. The above-referenced materials are considered hydrophobicwith respect to the contemplated environment of the invention. Surfacelayer materials include fluorosilicones and polyethylene glycol (PEG),polysaccharides, phospholipids, and combinations of the foregoing.

While heparin is preferred as the incorporated active material, agentspossibly suitable for incorporation include antithrobotics,anticoagulants, antibiotics, antiplatelet agents, thorombolytics,antiproliferatives, steroidal and non-steroidal antinflammatories,agents that inhibit hyperplasia and in particular restenosis, smoothmuscle cell inhibitors, growth factors, growth factor inhibitors, celladhesion inhibitors, cell adhesion promoters and drugs that may enhancethe formation of healthy neointimal tissue, including endothelial cellregeneration. The positive action may come from inhibiting particularcells (e.g., smooth muscle cells) or tissue formation (e.g.,fibromuscular tissue) while encouraging different cell migration (e.g.,endothelium) and tissue formation (neointimal tissue).

Suitable materials for fabricating the braided stent include stainlesssteel, tantalum, titanium alloys including nitinol (a nickel titanium,thermomemoried alloy material), and certain cobalt alloys includingcobalt-chromium-nickel alloys such as Elgiloy® and Phynox®. Furtherdetails concerning the fabrication and details of other aspects of thestents themselves may be gleaned from the above referenced U.S. Pat.Nos. 4,655,771 and 4,954,126 to Wallsten and U.S. Pat. No. 5,061,275 toWallsten et al, which are incorporated by reference herein.

Various combinations of polymer coating materials can be coordinatedwith biologically active species of interest to produce desired effectswhen coated on stents to be implanted in accordance with the invention.Loadings of therapeutic materials may vary. The mechanism ofincorporation of the biologically active species into the surfacecoating and egress mechanism depend both on the nature of the surfacecoating polymer and the material to be incorporated. The mechanism ofrelease also depends on the mode of incorporation. The material mayelute via interparticle paths or be administered via transport ordiffusion through the encapsulating material itself.

For the purposes of this specification, "elution" is defined as anyprocess of release that involves extraction or release by direct contactof the material with bodily fluids through the interparticle pathsconnected with the exterior of the coating. "Transport" or "diffusion"are defined to include a mechanism of release in which the materialreleased traverses through another material.

The desired release rate profile can be tailored by varying the coatingthickness, the radial distribution (layer to layer) of bioactivematerials, the mixing method, the amount of bioactive material, thecombination of different matrix polymer materials at different layers,and the crosslink density of the polymeric material. The crosslinkdensity is related to the amount of crosslinking which takes place andalso the relative tightness of the matrix created by the particularcrosslinking agent used. This, during the curing process, determines theamount of crosslinking and also the crosslink density of the polymermaterial. For bioactive materials released from the crosslinked matrix,such as heparin, a denser crosslink structure will result in a longerrelease time and reduced burst effect.

It will also be appreciated that an unmedicated silicone thin top layerprovides some advantage and additional control over drug elusion;however, in the case of heparin, for example, it has been found that atop coat or surface coating modified to further control the initialheparin release profile or to make the surface more non-thrombogenicpresents a distinct advantage.

BRIEF DESCRIPTION OF THE DRAWINGS

In the drawings, wherein like numerals designate like parts throughoutthe same:

FIG. 1 is a schematic flow diagram illustrating the steps of the processof the invention;

FIG. 2 represents a release profile for a multi-layer system showing thepercentage of heparin released over a two-week period;

FIG. 3 represents a release profile for a multi-layer system showing therelative release rate of heparin over a two-week period;

FIG. 4 illustrates a profile of release kinetics for different drugloadings at similar coating thicknesses illustrating the release ofheparin over a two-week period without associated means to provide along term non-thrombogenic surface thereafter;

FIG. 5 illustrates drug elution kinetics at a given loading of heparinover a two-week period at different coating thicknesses withoutassociated means to provide a long term non-thrombogenic surfacethereafter;

FIG. 6 illustrates the release kinetics for a given undercoat andtopcoat material varied according to thickness in which the percentageheparin in the undercoat and topcoats are kept constant;

FIG. 7 is a plot of heparin release kinetics in phosphate buffer systemat PH 7.4 with and without fluorosilicone (FSi) topcoat; and

FIG. 8 is another plot of heparin release kinetics in phosphate buffersystem in which a topcoat containing fluorosilicone (FSi) only iscompared with an FSi topcoat containing 16.7% imbedded heparin.

DETAILED DESCRIPTION

According to the present invention, the stent coatings incorporatingbiologically active materials for timed delivery in situ in a body lumenof interest are preferably sprayed in many thin layers from preparedcoating solutions or suspensions. The steps of the process areillustrated generally in FIG. 1. The coating solutions or suspensionsare prepared at 10 as will be described later. The desired amount ofcrosslinking agent (if any) is added to the suspension/solution as at 12and material is then agitated or stirred to produce a homogenous coatingcomposition at 14 which is thereafter transferred to an applicationcontainer or device which may be a container for spray painting at 16.Typical exemplary preparations of coating solutions that were used forheparin and dexamethasone appear next.

General Preparation of Heparin Undercoating Composition

Silicone was obtained as a polymer precursor in solvent (xylene)mixture. For example, a 35% solid silicone weight content in xylene wasprocured from Applied Silicone, Part #40,000. First, the silicone-xylenemixture was weighed. The solid silicone content was determined accordingto the vendor's analysis. Precalculated amounts of finely dividedheparin (2-6 microns) were added into the silicone, then tetrahydrofuron(THF) HPCL grade (Aldrich or EM) was added. For a 37.5% heparin coating,for example: W_(silicone) =5 g; solid percent=35%; W_(hep)=5×0.35×0.375/(0.625)=1.05 g. The amount of THF needed (44 ml) in thecoating solution was calculated by using the equation W_(silicone) solid/V_(THF) =0.04 for a 37.5% heparin coating solution). Finally, themanufacturer crosslinker solution was added by using Pasteur P-pipet.The amount of crosslinker added was formed to effect the release rateprofile. Typically, five drops of crosslinker solution were added foreach five grams of silicone-xylene mixture. The solution was stirred byusing the stirring rod until the suspension was homogenous andmilk-like. The coating solution was then transferred into a paint jar incondition for application by air brush.

General Preparation of Dexamethasone Undercoating Composition

Silicone (35% solution as above) was weighed into a beaker on a Metlerbalance. The weight of dexamethasone free alcohol or acetate form wascalculated by silicone weight multiplied by 0.35 and the desiredpercentage of dexamethasone (1 to 40%) and the required amount was thenweighed. Example: W_(silicone) =5 g; for a 10% dexamethasone coating,W_(dex) =5×0.35×0.1/0.9=0.194 g and THF needed in the coating solutioncalculated. W_(silicone) solid /V_(THF) =0.06 for a 10% dexamethasonecoating solution. Example: W_(silicone) =5 g; V_(THF) =5×0.35/0.06≈29ml. The dexamethasone was weighed in a beaker on an analytical balanceand half the total amount of THF was added. The solution was stirredwell to ensure full dissolution of the dexamethasone. The stirredDEX-THF solution was then transferred to the silicone container. Thebeaker was washed with the remaining THF and this was transferred to thesilicone container. The crosslinker was added by using a Pasteur pipet.Typically, five drops of crosslinker were used for five grams ofsilicone.

The application of the coating material to the stent was quite similarfor all of the materials and the same for the heparin and dexamethasonesuspensions prepared as in the above Examples. The suspension to beapplied was transferred to an application device, at 16 in FIG. 1.Typically a paint jar attached to an air brush, such as a Badger Model150, supplied with a source of pressurized air through a regulator(Norgren, 0-160 psi) was used. Once the brush hose was attached to thesource of compressed air downstream of the regulator, the air wasapplied. The pressure was adjusted to approximately 15-25 psi and thenozzle condition checked by depressing the trigger.

Any appropriate method can be used to secure the stent for spraying androtating fixtures were utilized successfully in the laboratory. Bothends of the relaxed stent were fastened to the fixture by two resilientretainers, commonly alligator clips, with the distance between the clipsadjusted so that the stent remained in a relaxed, unstretched condition.The rotor was then energized and the spin speed adjusted to the desiredcoating speed, nominally about 40 rpm.

With the stent rotating in a substantially horizontal plane, the spraynozzle was adjusted so that the distance from the nozzle to the stentwas about 2-4 inches and the composition was sprayed substantiallyhorizontally with the brush being directed along the stent from thedistal end of the stent to the proximal end and then from the proximalend to the distal end in a sweeping motion at a speed such that onespray cycle occurred in about three stent rotations. Typically a pauseof less than one minute, normally about one-half minute, elapsed betweenlayers. Of course, the number of coating layers did and will vary withthe particular application. For example, typical tie-layers as at 18 inFIG. 1, for a coating level of 3-4 mg of heparin per cm² of projectedarea, 20 cycles of coating application are required and about 30 ml ofsolution will be consumed for a 3.5 mm diameter by 14.5 cm long stent.

The rotation speed of the motor, of course, can be adjusted as can theviscosity of the composition and the flow rate of the spray nozzle asdesired to modify the layered structure. Generally, with the abovemixes, the best results have been obtained at rotational speeds in therange of 30-50 rpm and with a spray nozzle flow rate in the range of4-10 ml of coating composition per minute, depending on the stent size.It is contemplated that a more sophisticated, computer-controlledcoating apparatus will successfully automate the process demonstrated asfeasible in the laboratory.

Several applied layers make up what is called the undercoat as at 18. Inone process, additional upper undercoat layers, which may be of the sameor different composition with respect to bioactive material, the matrixpolymeric materials and crosslinking agent, for example, may be appliedas the top layer as at 20. The application of the top layer follows thesame coating procedure as the undercoat with the number and thickness oflayers being optional. Of course, the thickness of any layer can beadjusted by adjusting the speed of rotation of the stent and thespraying conditions. Generally, the total coating thickness iscontrolled by the number of spraying cycles or thin coats which make upthe total coat.

As shown at 22 in FIG. 1, the coated stent is thereafter subjected to acuring step in which the pre-polymer and crosslinking agents cooperateto produce a cured polymer matrix containing the biologically activespecies. The curing process involves evaporation of the solvent xylene,THF, etc. and the curing and crosslinking of the polymer. Certainsilicone materials can be cured at relatively low temperatures, (i.e.RT-50° C.) in what is known as a room temperature vulcanization (RTV)process. More typically, however, the curing process involves highertemperature curing materials and the coated stents are put into an ovenat approximately 90° C. or higher for approximately 16 hours. Thetemperature may be raised to as high as 150° C. for dexemethasanecontaining coated stents. Of course, the time and temperature may varywith particular silicones, crosslinkers and biologically active species.

Stents coated and cured in the manner described need to be sterilizedprior to packaging for future implantation. For sterilization, gammaradiation is a preferred method particularly for heparin containingcoatings; however, it has been found that stents coated and curedaccording to the process of the invention subjected to gammasterilization may be too slow to recover their original posture whendelivered to a vascular or other lumen site using a catheter unless apretreatment step as at 24 is first applied to the coated, cured stent.

The pretreatment step involves an argon plasma treatment of the coated,cured stents in the unconstrained configuration. In accordance with thisprocedure, the stents are placed in a chamber of a plasma surfacetreatment system such as a Plasma Science 350 (Himont/Plasma Science,Foster City, Calif.). The system is equipped with a reactor chamber andRF solid-state generator operating at 13.56 mHz and from 0-500 wattspower output and being equipped with a microprocessor controlled systemand a complete vacuum pump package. The reaction chamber contains anunimpeded work volume of 16.75 inches (42.55 cm) by 13.5 inches (34.3cm) by 17.5 inches (44.45 cm) in depth.

In the plasma process, unconstrained coated stents are placed in areactor chamber and the system is purged with nitrogen and a vacuumapplied to 20-50 mTorr. Thereafter, inert gas (argon, helium or mixtureof them) is admitted to the reaction chamber for the plasma treatment. Ahighly preferred method of operation consists of using argon gas,operating at a power range from 200 to 400 watts, a flow rate of 150-650standard ml per minute, which is equivalent to 100-450 mTorr, and anexposure time from 30 seconds to about 5 minutes. The stents can beremoved immediately after the plasma treatment or remain in the argonatmosphere for an additional period of time, typically five minutes.

After this, as shown at 26, the stents may be exposed to gammasterilization at 2.5-3.5 Mrad. The radiation may be carried out with thestent in either the radially non-constrained status--or in the radiallyconstrained status.

Preferably, however, the surface is modified prior to plasma treatmentor just prior to sterilization by one of several additional processingmethods of which some are described in relation to the followingexamples.

EXAMPLE 1

Fluorosilicone Surface Treatment of Eluting Heparin Coating

The undercoat of a stent was coated as multiple applied layers asdescribed above thereafter and cured as described at 22. The heparincontent of the undercoat was 37.5% and the coating thickness was about30-40μ. Fluorosilicone (FSi) spray solution was prepared at 30 from afluorosilicone suspension (Applied Silicone #40032) by weighing anamount of fluorosilicone suspension and adding tetrahydrofuran (THF)according to the relation equation of V_(THF) =1.2×the weight offluorosilicone suspension. The solution was stirred very well andspray-coated on the stent at 32 using the technique of the applicationof the undercoat process at 18 and the coated stents were cured at 90°C. for 16 hours. The coated stents are argon plasma treated prior togamma sterilization according to the procedures described above inaccordance with steps 22-26.

FIG. 7 is a plot of heparin release kinetics in phosphate buffer systemwith fluorosilicone topcoat and without any topcoat. The thickness ofthe topcoat is about 10-15μ. While it does not appear on the graph ofFIG. 7, it should be noted that the release rate for the coating withoutFSi is initially about 25 times higher than that with FSi, i.e., duringthe first 2 hours. This is, of course, clearly off the scale of thegraph. It is noteworthy, however, that the coating with the FSi toplayer or diffusion barrier does show a depressed initial release ratecombined with an enhanced elusion rate after the first day and throughthe first week up until about the tenth day. In addition, thefluorosilicone (FSi) topcoat, by virtue of the high electro-negativityof fluorination maintains non-thrombogenic surface qualities during andafter the elusion of the biologically active heparin species. Inaddition, because of the negative charges on the heparin itself, theelectro-negativity of the fluorosilicone topcoat may be, at least inpart, responsible for the modified heparin release kinetic profile.

FIG. 8 compares a plot of fluorosilicone (FSi) top coating containing16.7% imbedded heparin with one containing fluorosilicone (FSi) only. Anundercoating is identical to that utilized in FIG. 7 containing about37.5% heparin to a thickness of about 30-40 microns. These elutionkinetics are quite comparable with the heparin-free FSi top layergreatly reducing the initial burst of heparin release and otherwise theheparin in the FSi top layer imparts a slightly greater release over theperiod of the test.

EXAMPLE 2

Immobilization of Polyethylene Glycol (PEG) on Drug Eluting Undercoat

An undercoat was coated on a stent and cured at 22 as in Example 1. Thestent was then treated by argon gas plasma as at 24 and ammonium gasplasma at 40. The equipment and the process of argon gas plasmatreatment was as has been described above. The ammonium plasma treatmentwas implemented immediately after the argon gas plasma treatment, toaminate the surface of the coating. The ammonium flow rate was in therange of 100-700 cubic centimeter per minute (ccM) in preferably in therange of 500-600 ccM. The power output of radio frequency plasma was inthe range of 50-500 watts, preferably in ˜200 watts. The process timewas in the range of 30 sec-10 min, preferably ˜5 min.

Immediately after amination, the stents were immersed intoelectrophilically activated polyethylene glycol (PEG) solution at 42.PEG is known to be an inhibitor of protein absorption. Examples ofelectrophilically activated PEG are PEG nitrophenyl carbonates, PEGtrichlorophenyl carbonates, PEG tresylate, PEG glycidyl ether, PEGisocyanate, etc., optionally with one end terminated with methoxylgroup. Molecular weight of PEG ranged from about 1000-6000, and ispreferable about 3000. It has been observed that simple ammoniumamination will not generate large quantities of primary and secondaryamines on the elastomeric polymer surface (for example silicone).Instead, imine (>C═N--H), and other more oxidative nitro containinggroups will dominate the surface. It is generally necessary to addreductive agent such as NaBH₃ CN into the reaction media so that thefunctional group on PEG can react with imine and possibly othernitro-containing species on the surface, and therefore immobilize PEGonto the surface. The typical concentration of NaBH₃ CN is about 2mg/ml. Since PEG and its derivatives dissolve in water and many polarand aromatic solvents, the solvent used in the coating must be a solventfor PEG but not for the drug in the undercoat to prevent the possibleloss of the drug through leaching. In the case of eluting-heparincoating, a mixed solvent of formamide and methyl ethyl ketone (MEK) or amixed solvent of formamide and acetone are preferred solvents(preferably at ratios of 30 formamide: 70 MEK or acetone by volume),since they will not dissolve heparin. The concentration of PEG, thereaction time, the reaction temperature and the pH value depend on thekind of PEG employed. In the case of eluting heparin coating, 5% PEGtresylate in (30-70) Formamide/MEK was used successfully. The reactiontime was 3 hours at room temperature. PEG was then covalently bound tothe surface. Gamma radiation was then used for sterilization of thisembodiment as previously described.

With respect to the anticoagulant material heparin, the percentage inthe undercoat is nominally from about 30-50% and that of the topcoatfrom about 0-30% active material. The coating thickness ratio of thetopcoat to the undercoat varies from about 1:10 to 1:2 and is preferablyin the range of from about 1:6 to 1:3.

Suppressing the burst effect also enables a reduction in the drugloading or in other words, allows a reduction in the coating thickness,since the physician will give a bolus injection ofantiplatelet/anticoagulation drugs to the patient during the stentingprocess. As a result, the drug imbedded in the stent can be fully usedwithout waste. Tailoring the first day release, but maximizing secondday and third day release at the thinnest possible coating configurationwill reduce the acute or subacute thrombosis.

FIG. 4 depicts the general effect of drug loading for coatings ofsimilar thickness. The initial elution rate increases with the drugloading as shown in FIG. 5. The release rate also increases with thethickness of the coating at the same loading but tends to be inverselyproportional to the thickness of the topcoat as shown by the same drugloading and similar undercoat thickness in FIG. 6.

What is apparent from the data gathered to date, however, is that theprocess of the present invention enables the drug elution kinetics to becontrolled in a manner desired to meet the needs of the particular stentapplication. In a similar manner, stent coatings can be prepared using acombination of two or more drugs and the drug release sequence and ratecontrolled. For example, antiproliferation drugs may be combined in theundercoat and antiplatelet drugs in the topcoat. In this manner, theantiplatelet drugs, for example, heparin, will elute first followed byantiproliferation drugs to better enable safe encapsulation of theimplanted stent.

The heparin concentration measurement were made utilizing a standardcurve prepared by complexing azure A dye with dilute solutions ofheparin. Sixteen standards were used to compile the standard curve in awell-known manner.

For the elution test, the stents were immersed in a phosphate buffersolution at pH 7.4 in an incubator at approximately 37° C. Periodicsamplings of the solution were processed to determine the amount ofheparin eluted. After each sampling, each stent was placed inheparin-free buffer solution.

As stated above, while the allowable loading of the elastomeric materialwith heparin may vary, in the case of silicone materials heparin mayexceed 60% of the total weight of the layer. However, the loadinggenerally most advantageously used is in the range from about 10% to 45%of the total weight of the layer. In the case of dexamethasone, theloading may be as high as 50% or more of the total weight of the layerbut is preferably in the range of about 0.4% to 45%.

It will be appreciated that the mechanism of incorporation of thebiologically active species into a thin surface coating structureapplicable to a metal stent is an important aspect of the presentinvention. The need for relatively thick-walled polymer elution stentsor any membrane overlayers associated with many prior drug elutiondevices is obviated, as is the need for utilizing biodegradable orreabsorbable vehicles for carrying the biologically active species. Thetechnique clearly enables long-term delivery and minimizes interferencewith the independent mechanical or therapeutic benefits of the stentitself.

Coating materials are designed with a particular coating technique,coating/drug combination and drug infusion mechanism in mind.Consideration of the particular form and mechanism of release of thebiologically active species in the coating allow the technique toproduce superior results. In this manner, delivery of the biologicallyactive species from the coating structure can be tailored to accommodatea variety of applications.

Whereas the above examples depict coatings having two different drugloadings or percentages of biologically active material to be released,this is by no means limiting with respect to the invention and it iscontemplated that any number of layers and combinations of loadings canbe employed to achieve a desired release profile. For example, gradualgrading and change in the loading of the layers can be utilized inwhich, for example, higher loadings are used in the inner layers. Alsolayers can be used which have no drug loadings at all. For example, apulsatile heparin release system may be achieved by a coating in whichalternate layers containing heparin are sandwiched between unloadedlayers of silicone or other materials for a portion of the coating. Inother words, the invention allows untold numbers of combinations whichresult in a great deal of flexibility with respect to controlling therelease of biologically active materials with regard to an implantedstent. Each applied layer is typically from approximately 0.5 microns to15 microns in thickness. The total number of sprayed layers, of course,can vary widely, from less than 10 to more than 50 layers; commonly, 20to 40 layers are included. The total thickness of the coating can alsovary widely, but can generally be from about 10 to 200 microns.

Whereas the polymer of the coating may be any compatible biostableelastomeric material capable of being adhered to the stent material as athin layer, hydrophobic materials are preferred because it has beenfound that the release of the biologically active species can generallybe more predictably controlled with such materials. Preferred materialsinclude silicone rubber elastomers and biostable polyurethanesspecifically.

This invention has been described herein in considerable detail in orderto comply with the Patent Statutes and to provide those skilled in theart with the information needed to apply the novel principles and toconstruct and use embodiments of the example as required. However, it isto be understood that the invention can be carried out by specificallydifferent devices and that various modifications can be accomplishedwithout departing from the scope of the invention itself.

We claim:
 1. A medical device having at least a portion which isimplantable into the body of a patient, wherein at least a part of thedevice portion is metallic and at least part of the metallic deviceportion is covered with a coating for release of at least onebiologically active material, wherein said coating comprises anundercoat comprising a hydrophobic elastomeric material incorporating anamount of biologically active material therein for timed releasetherefrom, and wherein said coating further comprises a topcoat which atleast partially covers the undercoat, said topcoat comprising abiostable, non-thrombogenic material which provides long termnon-thromobogenicity to the device portion during and after release ofthe biologically active material, and wherein said topcoat issubstantially free of an elutable material.
 2. The device of claim 1wherein said biologically active material is heparin.
 3. The device ofclaim 2 wherein the non-thrombogenic material is selected from the groupconsisting of fluorosilicone, polyethylene glycol (PEG),polysaccharides, phospholipids and combinations thereof.
 4. The deviceof claim 3 wherein the non-thrombogenic material is fluorosilicone. 5.The device of claim 3 wherein the non-thrombogenic material ispolyethylene glycol (PEG).
 6. The device of claim 1 wherein the medicaldevice is an expandable stent.
 7. The device of claim 1 wherein thetopcoat consists of a polymer.
 8. The device of claim 6 wherein thestent comprises a tubular body having open ends and an open latticesidewall structure and wherein the coating conforms to said sidewallstructure in a manner that preserves said open lattice.
 9. A stent forimplantation in a vascular lumen comprising a tubular body having openends and a sidewall and a coating on at least a part of a surface ofsaid sidewall, said coating further comprising an undercoat comprising ahydrophobic elastomeric material incorporating an amount of finelydivided heparin therein for timed release therefrom, and wherein saidcoating further comprises a topcoat comprising an amount offluorosilicone which is capable of providing long termnon-thrombogenicity to the surface during and after release of thebiologically active material, wherein said topcoat at least partiallycovers the undercoat, and wherein said topcoat is substantially free ofan elutable material.
 10. The device of claim 9 wherein the sidewall isan open lattice structure and wherein the coating conforms to saidsidewall structure in a manner that preserves said open lattice.
 11. Astent for implantation in a vascular lumen comprising a tubular bodyhaving open ends and a sidewall and a coating on at least a part of thesurface of said sidewall, said coating further comprising an undercoatcomprising a hydrophobic elastomeric material incorporating an amount offinely divided heparin therein for timed release therefrom, and whereinsaid coating further comprises a topcoat comprising an amount ofpolyethylene glycol (PEG) which is capable of providing long termnon-thrombogenicity to the surface during and after release of thebiologically active material, wherein said topcoat at least partiallycovers the undercoat, and wherein said topcoat is substantially free ofan elutable material.
 12. The device of claim 11 wherein the sidewall isan open lattice structure and wherein the coating conforms to saidsidewall structure in a manner that preserves said open lattice.